Radiation emitting devices are generally known and used as imaging and as radiation therapy devices for the treatment of patients.
Collimators are used in a wide variety of equipment in which it is desired to permit only beams of radiation emanating along a particular path to pass beyond a selected point or plane. Collimators are frequently used in radiation imagers to ensure that only radiation beams emanating along a direct path from the known radiation source strike the detector, thereby minimizing detection of beams of scattered or secondary radiation. Collimator design affects the field-of-view, spatial resolution, and sensitivity of the imaging system.
Particularly in radiation imagers used for medical diagnostic analyses or for non-destructive evaluation procedures, it is important that only radiation emitted from a known source and passing along a direct path from that source through the subject under examination be detected and processed by the imaging equipment. If the detector is struck by undesired radiation, i.e., radiation passing along non-direct paths to the detector, such as rays that have been scattered or generated in secondary reactions in the object under examination, performance of the imaging system is degraded. Performance is degraded by lessened spatial resolution and lessened contrast resolution that result from the detection of the scattered or secondary radiation rays. Examples of imagers and collimators for such imagers are disclosed in U.S. Pat. Nos. 6,556,657; 6,507,642; 6,505,966; 6,396,902; 6,388,816; 6,377,661; and 6,271,524, all of which are incorporated herein by reference.
Collimators are positioned to substantially absorb the undesired radiation before it reaches the detector. Collimators are traditionally made of a material that has a relatively high atomic number, such as tungsten, placed so that radiation approaching the detector along a path other than one directly from the known radiation source strikes the body of the collimator and is absorbed before being able to strike the detector. In a typical detector system, the collimator includes barriers extending outwardly from the detector surface in the direction of the radiation source so as to form channels through which the radiation must pass in order to strike the detector surface.
Some radiation imaging systems, such as computed tomography (CT) systems used in medical diagnostic work, or such as industrial imaging devices, use a point (i.e. a relatively small, such as 1 mm in diameter or smaller) source of x-ray radiation to illuminate the subject under examination. The radiation passes through the subject and strikes a radiation detector positioned on the side of the subject opposite the radiation source. In a CT system, the radiation detector typically comprises a one-dimensional array of detector elements. Each detector element is disposed on a module, and the modules are typically arranged end to end along a curved surface to form a radiation detector arm. The distance to the center of the module, on any one of the separate modules is the same, i.e., each panel is at substantially the same radius from the radiation source. On any given module there is a difference from one end of the module to the other in the angle of incidence of the radiation beams arriving from the point source.
For example, in a common medical CT device, the detector is made up of a number of x-ray detector modules, each of which has dimensions of about 32 mm by 16 mm, positioned along a curved surface having a radius of about 1 meter from the radiation point source. Each detector module has about 16 separate detector elements about 32 mm long by 1 mm wide arranged in a one-dimensional array, with collimator plates situated between the elements and extending outwardly from the panel to a height above the surface of the panel of about 8 mm. As the conventional CT device uses only a one-dimensional array (i.e., the detector elements are aligned along only one row or axis), the collimator plates need only be placed along one axis, between each adjoining detector element. Even in an arrangement with a panel of sixteen 1 mm-wide detector elements adjoining one another (making the panel about 16 mm across), if the collimator plates extend perpendicularly to the detector surface, there can be significant “shadowing” of the detector element by the collimator plates toward the ends of the detector module. This shadowing results from some of the beams of incident radiation arriving along a path such that they strike the collimator before reaching the detector surface. Even in small arrays as mentioned above (i.e. detector panels about 16 mm across), when the source is about 1 meter from the panel with the panel positioned with respect to the point source so that a ray from the source strikes the middle of the panel at right angles, over 7.5% of the area of the end detector elements is shadowed by collimator plates that extend 8 mm vertically from the detector surface. Even shadowing of this extent can cause significant degradation in imager performance as it results in non-uniformity in the x-ray intensity and spectral distribution across the detector module. In the one-dimensional array, the collimator plates can be adjusted slightly from the vertical to compensate for this variance in the angle of incidence of the radiation from the point source.
Advanced CT technology (e.g., volumetric CT), however, makes use of two-dimensional arrays, i.e., arrays of detector elements that are arranged in rows and columns. The same is true of the precision required for industrial imagers. In such an array, a collimator must separate each detector element along both axes of the array. The radiation vectors from the point source to each detector on the array have different orientations, varying both in magnitude of the angle and direction of offset from the center of the array. Additionally, detector arrays larger than the one-dimensional array discussed above may be advantageously used in imaging applications. As the length of any one panel supporting detector elements increases, the problem of the collimator structure shadowing large areas of the detector surface become more important. In any system using a “point source” of radiation and flat panels, some of the radiation beams that are desired to be detected. i.e., ones emanating directly from the radiation source to the detector surface, strike the detector surface at some angle offset from vertical.
Gamma ray imaging is currently used in medicine to obtain 3D images of patients' internal organs. One such gamma ray imaging device is disclosed in U.S. Pat. No. 6,271,524, which is incorporated herein by reference. Positron Emission Tomography (PET) is a medical gamma ray imaging technique frequently used for this purpose. Prior to conducting the imaging procedure, a patient is given a radio-pharmaceutical, which contains a positron emitting substance and which is selectively accumulated in a region of interest. When a positron emitted by the radio-pharmaceutical encounters an electron, the electron-positron pair annihilates, emitting two gamma photons of 511 keV each, flying in opposite directions. The simultaneous detection of these gamma photons by two gamma detectors positioned opposite to each other, indicates that a positron has been emitted and annihilated inside an organ of a patient. The simultaneous attribution of 2D coordinates to each one of the photons allows for the determination of the photon's line of flight. The position of the annihilation is along this line. When a multitude of gamma photon pairs are detected and the information is processed using appropriate algorithms, electronic circuitry, software, etc., a 3D image of the organ under examination can be reconstructed.
In radiation therapy, the device generally includes a gantry which can be swivelled around a horizontal axis of rotation in the course of a therapeutic treatment. Two such devices are disclosed in U.S. Pat. Nos. 6,526,123 6,240,161, both of which are incorporated herein by reference. A linear accelerator is located in the gantry for generating a high energy radiation beam for therapy. This high energy radiation beam can be an electron beam or photon (X-ray) beam. During treatment, this radiation beam is trained on one zone of a patient lying in the isocenter of the gantry rotation. To control the radiation emitted toward an object, a beam shielding device, such as a plate arrangement or a collimator, is typically provided in the trajectory of the radiation beam between the radiation source and the object.
A collimator is a beam shielding device which can include multiple leaves, for example, a plurality of relatively thin plates or rods, typically arranged as opposing leaf pairs. The plates themselves are formed of a relatively dense and radiation impervious material and are generally independently positionable to delimit the radiation beam. The beam shielding device defines a field on the object to which a prescribed amount of radiation is to be delivered. The usual treatment field shape results in a three-dimensional treatment volume which includes segments of normal tissue, thereby limiting the dose that can be given to the tumor. The dose delivered to the tumor can be increased if the amount of normal tissue being irradiated is decreased and the dose delivered to the normal tissue is decreased. Avoidance of delivery of radiation to the organs surrounding and overlying the tumor determines the dosage that can be delivered to the tumor. Once an analysis is completed as to the intensity level of radiation at a particular region on the body, the beam shielding device settings must be chosen according to the output number of fields. Often, the application of a particular sequence of radiation requires a prohibitive amount of time to deliver, or which is physically impossible for the beam shielding device to achieve. As a result, to provide a realizable dosage, fewer intensity levels of radiation must be provided, and/or fewer radiation fields are used, thus the dose volume histograms are thereby degraded. While methods are known to address deliver dosage demands according to the intensity maps (See U.S. Pat. No. 5,663,999), such systems still cause a degradation of the dose volume histogram.
Various methods have been used to manufacture thicker collimators. One method is to cast the collimator. Several methods of casting are disclosed in U.S. Pat. No. 3,988,589, which is incorporated herein by reference. One casting method is to cast the collimator as a single unit using removable pins in the mold to provide holes in the collimator. This method of manufacture, while producing an operational collimator, is impractical since, due to high friction between the cast lead and the pins and the fact that some collimators are convergent or divergent (to allow enlarged or miniaturized image formation) relative to the radiation source, each of the pins used to create the holes must be removed individually. This process is time consuming and costly, especially when one realizes that some such collimators have 1000 or more such holes. Another casting method is to cast thick corrugated lead sheets and assemble them. This alternative also is unsatisfactory due to joint leakage (i.e. the epoxied joints are permeable to high energy radiation) and to too much distorting radiation reaching the receiver of the medical device. Still another casting method is to cast a plurality of modules that are press fitted or cemented together to form the collimator.
Several other methods for forming collimators are disclosed in U.S. Pat. No. 4,450,706, which is incorporated herein by reference. One method includes the dissolving metal by a chemical reagent to form a specific collimator shape. Another method includes wrapping radiation-absorbing foils around a large number of mandrels. Another method involves the formation of a plurality of collimator strips which are folded transversely to their longitudinal extension such that the flat portions of two adjacent strips engage each other, whereby the outwardly extending portions of these two adjacent strips extend in opposite directions to form a series of parallel channels. Still another method involves the use of strips that have been stamped into a shape and subsequently bonded together.
The casting methods described above for manufacturing a collimator can only be used to fabricate relatively simple collimators having high error tolerances in design. As technology has advanced, a need for more complex collimators has arisen wherein such collimators have very low error tolerances. One manufacturing method to address this problem is disclosed in U.S. Pat. No. 6,377,661, which is incorporated herein by reference. This patent discloses a collimator manufacturing process which includes the steps of generating a computer-aided-drawing (AutoCAD) drawing of a two-dimensional (2D) collimator based upon overall imager system parameters, generating a stereo-lithographic (STL) file or files corresponding to the AutoCAD drawing and to the chosen size, position and orientation of the focally aligned channels to be formed in the collimator, and interfacing the STL files with machining equipment to machine out the material to be removed from a solid slab (workpiece) of radiation-absorbing material, to form the plurality of focally aligned channels extending through the workpiece.
Another method for manufacturing a collimator is disclosed in United States Patent Publication No. 2003/0128813 published on Jul. 10, 2003 entitled “Devices, methods, and systems involving cast computed tomography collimators” and 2003/0128812 published on Jul. 10, 2003 entitled “Devices, methods, and systems involving cast collimators”, both of which are incorporated herein by reference. In this patent publication, a cast computed-tomography collimator is formed from a litho graphically-derived micro-machined metallic foil stack lamination mold. The mold has a stacked plurality of micro-machined metallic foil layers. The mold is filled with a first casting material to form a collimator.
Although these casting techniques have improved the quality of collimator production, the casting process still cannot meet certain tolerances that are needed for highly sensitive medical devices. In view of the prior art, there is a need for a manufacturing process for a collimator that is cost effective, not overly time consuming to manufacture, and which can produce a very precise collimator in a variety of shapes and sizes.